The present invention relates to radiation detection and, more specifically, to a radiation detector comprising a scintillator and a semiconductor drift photodetector wherein the scintillator decay time and drift time in the photodetector are matched in order to achieve the best signal-to-noise ratio.
For gamma-ray spectroscopy applications there are two primary detection systems from which one may choose, namely: the scintillator/photomultiplier tube (PMT) combination and germanium. Scintillator/PMT systems offer excellent efficiency but have relatively poor energy resolution. PMT""s are also bulky and fragile components that are vulnerable to damage, have low and non-uniform quantum efficiency, require a stabilized high voltage for operation, and are undesirably affected by magnetic fields. Consequently, there has been an ongoing need for suitable alternatives to replace either the PMT, or the entire scintillator/PMT system, for gamma-ray spectrometry.
Germanium semiconductor detectors are excellent in terms of their energy resolution capabilities. They are also essentially insensitive to magnetic fields, a problem which must be considered when using PMTs. However, Ge devices do have very significant disadvantages, including a requirement that they be cooled to temperatures below 100xc2x0 K for proper operation and for insuring adequate resistance to radiation damage.
There are areas of investigation aimed at finding solid state alternatives to replace the PMT in gamma-ray spectrometry. Silicon PIN semiconductor detectors have been studied for this application. The performance of large-area Si-PIN photo detectors in combination with scintillators is limited by their relatively high detector capacitance, which increases with the area of the detector, and a large leakage current which increases with the Si volume of the device. Enlarging the area of these devices above 200 mm2 results in severe performance penalty due to the increased electronic noise [James et al 1992]. In view of the maturity of Si technology, it is difficult to expect dramatic breakthroughs in achievable active area and depletion thickness.
Silicon avalanche photodiodes (APD""s) produced high expectations as a solid-state replacement for the PMT [McIntyre 1966, Webb et al 1974, Entine et al 1983, Iwanczyk 1991, James et al 1992]. However, even though good results continue to be reported [Moszynski 1998], reliability is a major obstacle, and availability of commercial large-area ( greater than 200 mm2) devices for this application are still not feasible.
A scintillation detector based on CsI(TI) scintillation crystals coupled with HgI2 photodetectors has produced excellent results for volumetric scintillators for high energy gamma-rays [Wang, Patt and Iwanczyk 1994]. However, there is currently a lack of availability of HgI2 crystals of sufficient quality and area to be suitable for photodetector fabrication.
A new photodetector device, the silicon drift photodetector (SDP), is derived from the silicon drift particle detector (SDD) [P. Rehak U.S. Pat. No. 4,688,067] which historically has been linked to the charge coupled device (CCD) and was first implemented for detection and tracking of high energy particles [W. Chen, H. Kraner, Z. Li, P. Rehak, E. Gatti, A. Longni, M. Sampietro, P. Holl, J. Kemmer, U. Faschingbauer, B. Schmitt, A. Woner and J. P. Wurm, IEEE Trans. Nucl. Sci. Vol. 39, No.4,1992,619].
SDD structures have also been used for x-ray spectroscopy [G. Bertuccio, A. Castoldi, A. Longoni, M. Sampietro and C. Gautheir, Nucl. Inst. and Meth. Phys. Res. A312 (1992) 613; J. S. Iwanczyk, B. E. Patt, et. al. xe2x80x9cSimulation and Modeling of a New Silicon Drift Chamber X-ray Detector Design for Synchrotron Radiation Applicationsxe2x80x9d, Nucl. Instr. and Meth. in Phys. Res. A380 (1996) 288-294; J. Segal, J. Plummer, B. E. Patt, J. S. Iwanczyk, and G. Vilkelis, xe2x80x9cA New Structure for Controlling Dark Current Due to Surface Generation in Drift Detectors,xe2x80x9d manuscript in preparation; J. Segal, C. Aw, J. Plummer, C. Kenney, S. Parker, B. E. Patt, J. S. Iwanczyk, and G. Vilkelis, xe2x80x9cA Vertical High Voltage Termination Structure for High-Resistivity Silicon Detectors,xe2x80x9d Submitted to IEEE Nucl. Science Symposium and Medical Imaging Conference, 1997 [2] J. S. Iwanczyk, B. E. Patt, C. R. Tull, C. Kenney, J. Segal, J. Bradley, B. Hedman, and K. Hodgson, xe2x80x9cLarge Area Silicon Drift Detectors for X-Raysxe2x80x94New Results,xe2x80x9d Submitted to the 1998 IEEE Nuclear Science Symposium, Toronto, Canada, Nov 12-14, 1998].
Recently SDD structures for detection of light (hereinafter referred to as Silicon Drift Photodetectors (SDPs)) have also begun to appear. These include structures described by Hartman [Hartman, R., Struder L., Kemmer J., Lechner P., Lorenz E., and Mirzoyan R. Nuclear Instruments and Methods in Physics Research A387: 250-254 (1997); Olschner [Olschner F. IEEE Trans. Nucl. Sci. NS-43(3):1407-1410, (1996)] and Fiorini [Fiorini C., Perotti F., and Labanti C., IEEEE Transactions on Nuclear Science, V45, No3: 483-486 (1998)]. However, none of the Silicon Drift Photodetectors or combination of scintillator with the Silicon Drift Photodetector described were optimized, as they have not recognized nor taken into account the criteria of the present invention. Thus it is desirable to provide a means for implementing radiation detectors using scintillators with semiconductor drift photodetectors wherein the components are specifically constructed to achieve the best signal-to-noise ratio.
Radiation imaging systems typically are used to generate images of the distribution of radiation either transmitted through an object or emitted by an object. These images can be used to determine the structure and function of internal organs. The radiations are not of themselves visible to the naked eye. In emission imaging (xe2x80x9cNuclear Medicinexe2x80x9d), radiation invisible to the naked eye is generated within an organ by radiopharmaceutical or other radiation bearing substance which passes through or in some cases is designed to accumulate within the organ.
Prior emission imaging applications include single photon planar imaging and Single Photon Emission Computed Tomography (SPECT) for imaging the structure or function of internal organs. Anger introduced one system which has remained largely unchanged since it was first described in the 1950""s (Anger, HO xe2x80x9cScintillation camera,xe2x80x9d Rev. Sci. Instr. 29, 27. 1958; Anger, HO xe2x80x9cScintillation camera with multichannel collimators,xe2x80x9d J. Nucl. Med. 5, p515-531. 1964). These Anger-type gamma-ray cameras employed in single-photon emission imaging applications typically have a shielded enclosure, preferably made from lead or similar materials. Incident gamma-rays pass through the parallel-hole collimator which xe2x80x9cfocusesxe2x80x9d the gamma-rays. The collimator limits the system sensitivity for typical medical imaging applications. Gama-rays received through the collimator enter a large scintillation crystal (typically NaI[TI], CsI[TI] or CsI[Na]), generating light photons which pass through an optical diffuser to neighboring photomultiplier tubes. The light photons are guided through the scintillator and the optical diffuser using reflectors along the sides of the scintillator and diffuser. The light photons strike an array of the photomultiplier tubes, each of which is between 1 inch and 2 inches in diameter, and signals pass to analog electronics which perform the position arithmetic and spectrometric functions. A device can be used to display the acquired images. Below the photomultiplier tubes is a position/pulse-height module (Webb S, In xe2x80x9cThe Physics of Medical Imaging,xe2x80x9d Adam Hilger, Bristol, England p161 1996).
A second type of emission imaging in Nuclear Medicine is dual photon imaging, such as Positron Emission Tomography (PET). Positron Emission Tomography systems typically are used to generate images of the distribution of positron emitting radiopharmaceuticals or other positron radiation bearing substances in the body. This is done by locating the origin of radiation caused by annihilation of the positrons. In PET, two coincident and oppositely traveling gamma-rays, each with energy of 511 keV, are produced by positron annihilation in the tissue of the organ being imaged.
The advantage of this type of imaging system is that the sensitivity can be significantly better than a collimator-based system because no collimator is required to determine the origin of the gamma-rays. Instead, the origin of the event is determined in the following way: The two gamma-rays resulting from the positron annihilation are detected in coincidence and the origin of the event is then known to lie along a line joining the collected gamma-rays. The precise location in space and the accurate timing of the arrival of the gamma-rays is critical so that they can be correlated in time with the annihilation event.
In order to perform the measurement at high spatial resolution, the typical PET system consists of rings of many small closely-packed detectors that surround a cross section of the body. The detector is used to measure a radiation which propagates from an organ through the rest of the body, the air, and any other medium between the body and the detector until it impinges on the detector. The radiation interacts with the detector to generate electrical signals representative of the detected radiation. The electrical signals can then be processed to generate an image on a video display device such as a computer monitor.
One method of generating electrical signals from detected radiation uses a scintillator and a photodetector. The scintillator is composed of a material that absorbs radiation of specified energy (e.g., gamma radiation of, for example, 511 keV) and converts it to visible light. The photodetector, in turn, converts the light emitted by the scintillator into electrical signals. This method is discussed in Chapters 8-10 of a book titled xe2x80x9cRadiation Detection and Measurement, 2nd Editionxe2x80x9d by Glenn F. Knoll, published in 1989 by John Wiley and Sons, Inc.
In order to accurately locate an event, the detector must completely absorb the 511 keV gamma-ray energy in a small volume. If absorption is only partial, the gamma-ray can deposit energy in more than one crystal, making the location ambiguous. If this happens, the events are usually discarded, resulting in lower instrument sensitivity.
Typically, when taking a PET image, there are millions of 511 keV photons striking the PET detectors per second, and it is the task of the PET system to identify which pairs of photons correspond to the same positron annihilation (this defines a ray or line of response through the body along which the annihilation occurred). The selection is done by choosing only those events that occur simultaneously (typically less than 12 nanoseconds). Thus, the detector must be dense, with a high effective atomic number, and have a very fast signal.
Inorganic scintillation detectors have these characteristics and are typically coupled to photomultiplier tubes (PMTs) that convert scintillation light from the 511 keV photon absorption into useful, fast signals for the PET systems.
Early PET systems had relatively large detectors discretely coupled to PMTs, providing a simple high performance solution for low resolution PET systems with relatively few detectors per ring [Hoffman E J, et al. J Nucl Med. 17:493-502, (1976)].
A more recent solution that has dominated PET detectors for several years, is a 2-D matrix of scintillators coupled to a 2 by 2 array of PMTs [Casey M E, et al. IEEE Trans Nucl Sci, NS-33:460-463, (1986); Cherry S R, et al. IEEE Trans. Nucl. Sci. NS-42:1064-1068, (1995)]. This solution was attractive in that it cost less (fewer and less expensive PMTs), and it produced better spatial and energy resolution with better coincidence timing.
A number of devices were built to avoid the use of the PMT entirely. One of the earliest candidates, the multiwire proportional chamber [Townsend D, et al. IEEE Trans. Nucl. Sci. NS-27(1):463-70, (1980)], had very high spatial resolution, and was less expensive than other types of PET systems. Unfortunately, the sensitivity for 511 keV photons was too low, and attempts to improve the sensitivity led to very poor coincidence timing.
Recently, liquid Xenon multiwire drift chambers have been considered for PET [Chepel VY, et al. Nucl. Instr. Meth. A367:58-61, (1995)], and while the performance characteristics in terms of spatial and timing resolution have been impressive, the need for cryogenics and the low stopping power of the liquid xenon for the 511 keV photon will likely limit this technology""s role in PET.
Another technique uses a barium fluoride scintillation crystal, whose UV photons are in turn detected and positioned by a wire chamber with TMAE fill gas [Suckling, J, et al. Nucl. Instr. Meth. A310:465-70, (1991)]. Although various aspects of this technology have excellent performance, barium fluoride has low stopping power for 511 keV gamma-rays.
The use of solid state detectors with scintillators has been attempted by a number of groups. The earliest attempts used individual solid state detectors, attached to a matrix of scintillation crystals in turn coupled to a single PMT for energy measurement and timing, to locate the correct crystal of interaction.
In the earliest work, solid state mercuric iodide photodetectors were investigated for this purpose [Barton J B, et al. IEEE Trans. Nucl. Sci. NS-30:671-675, (1983); Dahlbom M, et al. IEEE Trans. Nucl. Sci. NS-32:533-537, (1985)], and this work was soon followed by employing silicon PIN diodes [Derenzo S E, et al. IEEE Trans. Nucl. Sci. NS-30(1):665-70, (1983); Moses W W, et al. IEEE Trans. Nucl. Sci. NS-40(4):1036-40, (1993); Moses W W, et al. Nucl. Instr. Meth. A352:189-94, (1994); Moses W W, et al. IEEEE Trans. Nucl. Sci. NS-42(4):1085-9, (1995)]. This concept was generally thought to add complexity to the PET system with only modest improvement in performance. One drawback of these devices was that the solid state component lacked the timing capability needed for PET. Thus they were used only for localization in space, and a PMT had to be used in conjunction with the solid state detector to obtain the timing component of the measurement.
Some efforts have been placed on development of solid state detectors capable of providing both the spatial and timing capabilities needed. The most promising of this type of solid state device for PET had been the Avalanche Photodiode (APD). However, to date, its success has been moderate. Lecomte et al. [Lecomte R, et al. IEEE Trans. Nucl. Sci., NS-32(1):482-6, (1985); Lecomte R, et al. IEEE Trans. Nucl. Sci. NS-43(3): 1952-1957, (1996)] have been developing APD-based PET systems since 1985.
One such device employed xe2x80x9creach throughxe2x80x9d type APDs, using one APD per bismuth germanate (BGO) crystal.
Early timing measurements gave FWHM resolutions of about 30 ns for APD-BGO in conjunction with a plastic scintillator [Melcher C L, et al. IEEE Trans. Nucl. Sci. NS-39(4):502-5, (1992)], which means that an APD-APD system would have a resolving time on the order of 45-50 ns, compared to the less than 12 ns timing window for PMTs. Eventually, this APD-APD system gave 20 ns FWHM timing resolution with 20 to 40 ns timing windows. The poor timing is due to low light yield from BGO, which is 10-15% of NaI(TI), and is also due to the long decay constant (300 ns).
Recently, SDP structures were described for this application by [Avset B S, et al. Nucl. Instr. Meth. A288:131-136, (1990) and Olschner F. IEEE Trans. Nucl. Sci. NS-43(3):1407-1410, (1996)]. However, these SDPs do not address the PET application and do not have the necessary characteristics in terms of charge collection and timing to be useful for a PET application.
These prior SDP designs focused on the reduction of electronic noise by minimizing device capacitance at the expense of the electric field. The weak electric field results in a significant increase in charge collection time, which is unacceptable for PET timing requirements. One example is the structure described in [Olschner F. IEEE Trans. Nucl. Sci. NS-43(3):1407-1410, (1996)], where the photosensitive surface is biased only from strip side by means of a punch-through effect. The weak electric fields and long transit times for electrons lead to very poor timing characteristics, and the devices reported to date have had no practical use for timing applications such as positron emission coincidence. Instead, they were intended to be applied as a simple position sensing device (not coincidence sensing) used in conjunction with PMTs.
Lutetium orthosilicate (LSO) is a new scintillator [Holl I, et al. IEEE Trans. Nucl. Sci. NS-42(4):351-6, (1995)] that almost matches the stopping power of BGO, but has more light output than BGO (75% of NaI(Tl)) and a shorter decay constant (12 and 42 ns). LSO has been shown to give timing resolutions on the order of 2.6 ns [Holl I, et al. IEEE Trans. Nucl. Sci. NS-42(4):351-6, (1995); Schmelz C, et al. IEEE Trans. Nucl. Sci. NS-42(4):1080-4, (1995)], and is a very promising new development for the advancement of PET.
Radiation detectors according to one embodiment of the invention are implemented using scintillators combined with a semiconductor drift photodetectors wherein the components are specifically constructed in terms of their geometry, dimensions, and arrangement so that the scintillator decay time and drift time in the photodetector pairs are matched in order to achieve the best signal-to-noise ratio. Thus, the radiation detector of the present invention may include a scintillator having an exit window, the cross section of the exit window substantially matched to the cross section of an entrance window of an optically-coupled silicon drift photodetector; the silicon drift photodetector having dimensions and an electric field distribution creating a time spread of a signal generated in the silicon drift photodetector substantially matching a decay time of the scintillator. In addition, the detector may include electronics for amplification of electrical signals produced by the silicon drift photodetector, the amplification having a shaping time optimized with respect to the decay time of the scintillator and time spread of the signal in the silicon drift photodetector to substantially maximize the ratio of the signal to the electronic noise.